Radiography is a long known medical diagnostic imaging technique.
In a conventional radiography system, an X-ray source is actuated to direct a divergent area beam of X-rays through a patient. A cassette containing an X-ray sensitive phosphor screen and light and X-ray sensitive film is positioned in the X-ray path on the side of the patient opposite the source. X-radiation passing through the patient's body is attenuated to produce a shadow image of a portion of the patient through which the X-rays pass.
More recently, digital radiographic techniques and systems have been developed. In digital radiography, the source directs X-radiation through a patient's body to a detector assembly located in the beam path beyond the patient. The detector produces electrical signals defining the radiation pattern emergent from the patient. These signals are then processed to yield a visual display of the image.
The detector assembly includes an elongated array of individual discrete detector elements. Each detector element responds to incident X-radiation to produce an analog electrical charge signal indicative of such radiation. These analog electrical signals represent the radiation pattern or image emergent from the patient's body and incident on the detector array.
The analog signals are sampled and processed by imaging circuitry, primarily to improve their signal to noise ratio, and are subsequently digitized.
The digital signals are fed to a digital data processing unit (DPU). The data processing unit records and/or processes and enhances the digital data.
A display unit responds to appropriate digital data representing the image to convert the digital information back into analog form and to produce a visual display of the patient's internal body structure derived from the acquired image pattern of radiation.
The display unit can be coupled directly to the digital data processing unit for substantially real time imaging, or can be fed stored digital data from digital storage means such as tapes or disks representing patient images produced from earlier studies.
Digital radiography includes techniques in which a thin spread beam of X-radiation is used. In practice of this technique, often called "scan (or slit) projection radiography" (SPR), the spread beam is scanned across the patient, or the patient is movably interposed between the spread beam X-ray source and the detector assembly, the detector being maintained in continuous alignment with the beam. The relative movement effected between the source-detector arrangement and the patient's body scans a large portion of the body.
Discrete element detectors have been proposed comprising a single line of detector elements. Other proposals have included rectangular detector arrays of square detector elements.
Details of certain aspects of digital radiography systems such as described here are set forth in the following publications, hereby expressely incorporated by reference: Mattson, R. A., et al, "Design and Physical Characteristics of a Digital Chest Unit", S.P.I.E. Volume 314, Digital Radiography (1981).
Arnold, B. A. et al "Digital Radiography: An Overview" Proc. of S.P.I.E. Volume 273, March 1981;
Kruger, R. A. et al "A Digital Video Image Processor for Real Time X-Ray Subtraction Imaging" Optical Engineering Volume 17, No. 6 (1978);
U.S. patent application Ser. No. 444,605, filed Nov. 26, 1982 by Gary L. Barnes and entitled "Split Energy Level Radiation Detection";
U.S. patent application No. 542,384, filed Oct. 17, 1983 by Mattson, et al entitled "Improving Signal Characteristics in Digital Scan Projection Radiography".
An alternate proposal to the detector element array described above is a detector array consisting of two side by side vertical columns of square detector elements. One of the columns, however, is slightly vertically displaced, or offset, with respect to the other by a distance equal to one half the height of a single detector element. Such a configuration is described in the above incorporated application by Barnes.
It has also been proposed, where the detector array comprises a rectangular array of square detector elements, to improve the signal to noise ratio of the information developed by the detector, by the use of time delay and integrate (TDI) circuitry. An embodiment of such a proposed system is described in U.S. Pat. No. 4,383,327, issued May 10, 1983 to Kruger, which is hereby incorporated by reference. Such proposed TDI systems employ sampling at regular intervals of detector motion, and motion-synchronous shifting and adding of individual detector-produced analog charge signals. In such systems, the TDI circuitry can be integral with the detector elements.
Important advantages of scanning slit radiography are excellent scatter rejection and compatibility with digital image sensors. A significant disadvantage of such systems is the requirement of heavy X-ray tube loading that results from inefficient utilization of the X-ray output. This inefficiency arises from the aperture width which defines the spread beam subtending only a small solid angle at the focal spot of the X-ray tube.
To alleviate this disadvantage, it has been proposed to use a spread beam which is as thick as possible without unduly compromising the inherent good scatter rejection of such systems employing thin beams. See the above incorporated patent application by Mattson et al. As the spread beam is widened, however, the use of a rectangular detector array becomes more difficult. Such difficulty arises because, as the spread beam is thickened, more detector elements are needed, data rates become high and accordingly more difficult to handle, and TDI techniques must be used for shifting and adding data synchronously with scan motion so that data pertaining to each image portion, or "pixel" are properly superimposed to avoid image blurring.
Where time delay and integrate shifting and adding circuitry is employed, the detector element output signals are sampled at successive increments of detector movement equal to the length of a side of a single detector element. For reasons explained in more detail below, the spatial resolution of a rectangular detector array when used with TDI as described above is poorer than the maximum inherently obtainable resolution.
To facilitate understanding of both the prior art and the present invention, certain information and definitions relating to imaging optics are useful.
The ability of any optical element or system to resolve images is often described in terms of its "modulation transfer function" (MTF). Normally, the ability of an optical system to resolve a portion of an image decreases as the fineness of detail of the image portion (the number of lines per unit distance) increases. The number of lines per unit distance is frequently expressed as "line pairs per millimeter", and is known as the "spatial frequency" of the image portion of interest. The degradation of resolution as detail increases is manifested as a reduction in the contrast between the light and dark areas of the image portion. MTF is the function of contrast ratio versus the spatial frequency.
A rectangular detector element has an MTF in each of the x and y co-ordinates of its energy receiving face. In a square detector element MTF.sub.x =MTF.sub.y, and both functions can be demonstrated to be represented by the expression sinc (pf), where p is the length of one side of the square element receiving face, f is the spatial frequency sought to be imaged, and the sinc function is defined as sinc x=sin (.pi.x)/(.pi.x)
According to the above relationship, the x and y MTF's each are first reduced to 0 when the spatial frequency increases to f=1/p. This first zero is generally taken as representing the maximum spatial frequency (detail) which a square detector element can reliably image.
This phenonemon is one limiting factor on the resolving capability of any square detector, and is dependent upon its size, or "aperture". This parameter is referred to as the "aperture cutoff frequency".
A detector element is also limited by another resolution constraint known as the "Nyquist frequency". The Nyquist frequency is a spatial frequency above which the detector element cannot resolve separate lines. Rather than being a function of the detector size, however, the Nyquist frequency is related to the incremental distance at which successive samplings of the detector element output signal occur. The Nyquist cutoff frequency is relevant since the use of TDI circuitry requires repeated detector output samplings.
Where a row of square detectors, extending in the x coordinate in a rectangular array, is sampled once for each successive element width increment, (sampling distance) (as in the prior art) it can be shown, as set forth in publications referenced below, that the Nyquist frequency is only 1/(2p), along both coordinate axes. Therefore, in such a rectangular array, sampled as described, the Nyquist frequency is twice as limiting to resolution as is the aperture cutoff. Thus, the spatial frequency at which the resolving capability disappears under the Nyquist criterion is only half the frequency at which the resolving capability disappears under the aperture cutoff frequency criterion.
This means that, where a moving rectangular array of square elements is employed, and a row of elements is sampled only at successive increments of one detector width, the spatial resolution of such a detector is poorer than the maximum obtainable, as dictated by the detector element size, or aperture, under the aperture cutoff criterion. Also, aliasing artifacts will be present in an image derived from such a detector.
The Nyquist criterion is also applicable in the y coordinate of a rectangular array. In the y direction, the equivalent sampling distance between adjancent rows of square elements is the length p of one side of an element.
The following publications are hereby expressely incorporated by reference, for the assistance of those not conversant with this art, which are explanatory of the theory relating to these conclusions:
Sones, R. A., et al "A Method to Measure the MTF of Digital X-ray Systems" Medical Physics 11(2), March/April 1984, pages 166-171; Giger, M. L. et al "Investigation of Basic Imaging Properties in Digital Radiography: Modulation Transfer Function" Medical Physics 11(3) May/June 1984, at pages 287-295.
It has been seen that, where a rectangular array of square elements is sampled only once for each successive element width of motion, the system fails to take full advantage of the resolving power of the elements as dictated by their size.
The same conclusion applies to resolution in the y coordinate of the rectangular array. This is because the effective sampling distance between adjacent rows of the arrays is defined as p, the same as in the x direction.
Therefore, a rectangular array, sampled at incremental distance p fails to take full advantage of its inherent resolving power, in either the x or y coordinate.
It has been proposed that the use of an offset array, having only two columns, might be used to improve (reduce) the effective sampling distance increment in the y direction. This, however, is only a partial improvement, since it does not afford any reduction in the sampling distance in the x coordinate.
Moreover, since such detectors have only two columns of elements, and the detector array scans perpendicular to the columns, such an array has not been used in conjunction with TDI imaging circuitry. Thus, the signal-enhancing benefits of such circuitry have not been useful in conjunction with any known staggered, or offset, array.
It is an object of this invention to take maximum advantage of the resolving capabilities of the individual detector elements of the detector array, as defined by the aperture cutoff criterion, by eliminating the Nyquist frequency restriction in both x and y coordinates, while maintaining the full benefits of the use of time delay and integrate circuitry for enhancing the signal to noise ratio of the data acquired by the detector array.